Tissue engineering is a rapidly growing field encompassing a number of technologies aimed at replacing or restoring tissue and organ function. The key objective in tissue engineering is the regeneration of a defective tissue through the use of materials that can integrate into the existing tissue so as to restore normal tissue function. Tissue engineering, therefore, demands materials that can support cell over-growth, in-growth, or encapsulation and, in many cases, nerve regeneration.
Various crosslinkers have been used to crosslink biopolymer scaffolds, such as collagen scaffolds, in diverse tissue engineering fields [12-15]. Collagen in the body makes stabilization of collagen-based biomaterials and chemical cross-linking methods necessary to give materials that maintain the desired mechanical properties and stability during the desired implantation period [16]. Crosslinking methods can be divided into two general methodologies based on the crosslinker chemistry [16]. One crosslinking methodology makes use of bifunctional reagents, which can be used to bond amine groups of lysine or hydroxylysine by monomeric or oligomeric crosslinks. Based on the use of bifunctional reagents for crosslinking, glutaraldehyde (GA) has generally been applied for the crosslinking of collagen-based materials [17]. The use of hexamethylene diisocyanate (HMDIC) as a cross-linking agent was introduced by Chvapil et al [18]. GA cross-linking involves the formation of short (branched) aliphatic chains and pyridinium compounds [19, 20], while in HMDIC cross-linking aliphatic chains containing urea bonds are introduced between two adjacent amine groups [21]. Both GA and HMDIC cross-linking may lead to the presence of unreacted functional groups (probably aldehyde or amine groups after hydrolysis of isocyanate groups) in the collagen matrix, which can result in a cytotoxic reaction upon degradation of the collagen. Furthermore, it has been reported that GA cross-linked collagen-based biomaterials releases toxic GA (related) molecules from the biomaterial, which may result from unreacted GA present in the samples or from hydrolytic or enzymatic degradation products. This may also contribute to the cytotoxic reactions elicited by these materials both in vitro and in vivo [22. 23].
The GA crosslinkers has been used to bridge amine groups of lysine or hydroxylysine residues of collagen polypeptide chains. However, one major disadvantage of these cross-linking agents is the potential toxic effect of residual molecules when the biomaterial is exposed to biological environments. e.g., during in vivo degradation.
A second crosslinking methodology makes use of amide type crosslinkers. They could be formed by activation of the carboxylic acid groups of glutamic and aspartic acid residues followed by reaction of these activated carboxylic acid groups with amine groups of another polypeptide chain [24]. Cross-linking methods based on the concept of cross-linking by activation of carboxylic acid groups have been developed. The use of cyanamide for cross-linking of reconstituted collagen was first reported by Weadock et al [25]. However carbodiimide type crosslinkers, especially 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide hydrochloride (EDC) and N-hydroxysuccinimide NHS, offer the main advantage of lower toxicity and better compatibility over other crosslinkers [26]. The acyl azide activation method was used for cross-linking of pericardium. Using these methods, direct cross-linking of the polypeptide chains occurs, resulting in the formation of amide-type crosslinks [27].
In principal, no unreacted groups will be left in the material during crosslinking provided that reagents used for the activation of the carboxylic acid groups are easily removed. Cross-linking of collagen-based biomaterials using these methods resulted in materials with a similar resistance against degradation by bacterial collagenase compared with GA cross-linked materials. The influence of N-hydroxysuccinimide (NHS) on the activation of the carboxylic acid groups and subsequent cross-linking of the collagen material was studied [16].
The cornea is a transparent, avascular tissue, the structure of which allows it to serve as both a barrier to the outside environment and as an optical pathway. Vision loss due to corneal disease or trauma affects over 10 million individuals worldwide. For many, although treatable by corneal transplantation, donor tissue demand exceeds supply, especially in the developing countries [1-3]. While corneal substitutes have been proposed, to date, the only substitutes clinically tested in humans have been fully synthetic keratoprostheses (KPros). Although improving, complications with keratoprostheses, including retroprosthetic membrane formation, calcification, infection, and glaucoma, have limited their use to cases not treatable by human donor grafting [4]. Prostheses therefore do not alleviate the primary need for human donor corneas, especially in the developing world where the shortage of human donor corneas is acute.
An alternative approach is to enhance the inherent regenerative capacity of the human cornea to restore healthy, viable tissue. Tissue-engineered mimics of the extracellular matrix (ECM) have been proposed as scaffolds for endogenous tissue regeneration. In this regard, a range of biomimetic corneal substitutes have been developed, comprising either crosslinked medical grade porcine or recombinant human collagen [5] or hybrid collagen-synthetic [6-8] materials. These materials have provided robust, implantable, cornea-shaped scaffolds.
A simple biomimetic corneal substitute based on human collagen crosslinked with EDC (1-Ethyl-3-(3-dimethyl aminopropyl) carbodiimide hydrochloride) has been previously reported. This simple but biointeractive corneal substitute has been successfully tested in pig models, showing regeneration of corneal cells and nerves [5]. Although EDC was used successfully in previous experiments, the gelation time of collagen hydrogel crosslinked EDC was very short, making it very difficult to fabricate hydrogels. As a result, the fabrication process must be performed at cold temperature, preferably at 0-4° C., at which temperature the gelation may still be too quick for facilitating fabrication of hydrogels and biopolymers for various uses. Additionally, a short gelation time makes it difficult to produce hydrogels and biopolymers that incorporate corneal stem or progenitor cells using EDC as the crosslinker, since the collagen solution must be mixed with corneal fibroblasts before gelation of the collagen.
There remains a need for an alternative to EDC as a crosslinker in fabricating hydrogels. In particular, a method that would permit collagen, or another suitable biopolymer, to gel slowly at room temperature slowly would be particularly useful for producing hydrogels useful in various medical applications, including ophthalmic devices, such as, for example, corneal substitutes and corneal implants.
This background information is provided for the purpose of making known information believed by the applicant to be of possible relevance to the present invention. No admission is necessarily intended, nor should be construed, that any of the preceding information constitutes prior art against the present invention.